Magnetic Resonance Imaging (MRI) has become a widely accepted and commercially available technique for obtaining digitized visual images representing the internal structure of objects (such as the human body) having substantial populations of atomic nuclei that are susceptible to nuclear magnetic resonance (NMR) phenomena. In MRI nuclei in a body to be imaged are polarized by imposing a strong main magnetic field H.sub.0 on the nuclei. Selected nuclei are excited by imposing a radio frequency (RF) signal at a particular NMR frequency. By spatially distributing the localized magnetic fields, and then suitably analyzing the resulting RF responses from the nuclei, a map or image of relative NMR responses as a function of the location of the nuclei can be determined. Following a Fourier analysis, data representing the NMR responses in space can be displayed on a CRT.
As shown in FIG. 1, the NMR imaging system typically includes a magnet 10 to impose the static magnetic field, gradient coils 14 for imposing spatially distributed magnetic fields along three orthogonal coordinates, and RF coils 15 and 16 to transmit and receive RF signals to and from the selected nuclei. The NMR signal received by the coil 16 is transmitted to a computer 19 which processes the data into an image displayed on display 24. The magnetic resonance image is composed of picture elements called "pixels." The intensity of a pixel is proportional to the NMR signal intensity of the contents of a corresponding volume element or "voxel" of the object being imaged. The computer 19 also controls the operation of RF coils 15 and 16 and gradient coils 14 through the RF amplifier/detector 21 and 22 and gradient amplifiers 20, respectively.
Only nuclei with odd number of protons and/or neutrons have a magnetic moment and thus are susceptible to NMR phenomena. In MRI, a strong static magnetic field is employed to align nuclei, generating a gross magnetization vector aligned in parallel to the main magnetic field at equilibrium. A second magnetic field, applied transverse to the first field as is a single RF pulse, pumps energy into the nuclei, which causes the gross magnetization vector to flip by, for example, 90.degree.. After this excitation, the nuclei precess and gradually relax back into alignment with the static field. As the nuclei precess and relax, they will induce a weak but detectable electrical energy in the surrounding coils that is known as free induction decay (FID). These FID signals (and/or magnetic gradient-refocused field echoes thereof), collectively referred to herein as MR signals, are analyzed by a computer to produce images of the nuclei in space.
An operation whereby the various coils produces RF excitation pulses and gradient fields to result in and acquire an MR signal is called an MRI acquisition sequence. A graphical representation of an example MRI acquisition sequence used for 3-D MRI is shown in FIG. 2. In this example. the particular timing of applied pulses and fields is known as a field-echo sequence since the MR signals appear as gradient-refocused field echoes. First, a gradient field, G.sub.slice, is superimposed along the main field to sensitize a slab of nuclei in the body to be imaged to a particular RF resonance frequency. An RF excitation field or nutation pulse, .theta., is then applied at the particular frequency to tip the magnetization away from equilibrium. Thereafter, pulsed magnetic gradient fields of changing magnitudes, G.sub.pe and G.sub.slice, are used to phase encode the nuclei by inducing a temporary frequency difference, and hence phase differences, between nuclei in different locations along a specific direction within the slab. At the same time, another pulsed magnetic gradient field, G.sub.ro, is applied perpendicular to the direction of G.sub.pe, in a readout (ro) direction that first de-phases and then rephases the precessing nuclei--which results in producing field-echo MR signal S. The time from the center of nutating pulse, .theta., to the center of the field-echo MR signal is designated as the echo time, TE, and the entire pulse sequence duration is designated as TR.
Essentially, the applied gradient field, G.sub.ro, frequency encodes the selected slab of nuclei in the readout direction. The resultant MR signal, S, (also called "raw data" or "k-space data") is then read and analyzed by Fourier analysis. A frequency domain plot of that analysis is then scaled to render information about the nuclei population in Fourier space (also referred to as the image domain), which corresponds to an X-Y-Z position.
A magnetization vector can be decomposed into longitudinal and transverse components in reference to the main B.sub.0 field. Conventionally, the longitudinal component is defined as parallel to the B.sub.0 field and the transverse component is defined as perpendicular to B.sub.0. Once the magnetic vectors are disturbed from their equilibrium, processes known as "relaxation" cause the longitudinal component to recover to an equilibrium magnitude, M.sub.0, in alignment with the background B.sub.0 field, and the transverse component to decay. These relaxation processes are termed the "spin-lattice relaxation" and the "spin-spin relaxation" and are characterized by exponentials whose defined time constants are labeled as T.sub.1 and T.sub.2, respectively. In addition to T.sub.2 relaxation, inhomogeneities in magnetic field cause the transverse component to further decay. An "apparent relaxation" time constant, T.sub.2 *, is therefore defined as characterizing transverse signal decay due to both spin-spin relaxation and the presence of B.sub.0 field inhomogeneities.
The NMR frequency and the main B.sub.0 field are related by the Larmor relationship. This relationship states that the angular frequency, .omega..sub.0, of the precession of the nuclei is the product of the magnetic field, B.sub.0, and the so-called magnetogyric ratio, .gamma., a fundamental physical constant for each nuclear species: EQU .omega..sub.0 =B.sub.0 .multidot..gamma.(1-.sigma.)
where .sigma. is a shielding factor representing the chemical environment around the nuclei, commonly referred to as the "chemical shift."
The RF spin-nutating pulse will, of course, tip more than one species of the target isotope in a particular area. After being tipped away from equilibrium, each species of nuclei will begin to precess at their own characteristic speed. The phase of the precessing nuclei species will gradually differ (de-phase) as a result of such parameters as the physical or chemical environment in which the nuclei are located. Nuclei in fat, for example, precess at a different rate than do nuclei in water due to the effects of chemical shift. In addition, inhomogeneities in the magnetic field also contribute to de-phasing of the nutated precessing nuclei.
Since hydrogen nuclei have a readily discernible NMR signal and are the most abundant isotope of the human body, human MRI primarily images the NMR signal from the hydrogen nuclei. Water and fat are the main tissue components containing hydrogen nuclei.
In addition to using the frequency information content of an MR signal to generate images, the phase of an MR signal in the frequency domain can be utilized to provide information indicative of some physical quantity. For example, depending on the type of pulse sequence used, the MR phase can be used to differentiate between water and fat. It can also represent a main B.sub.0 field inhomogeneity or can be proportional to the velocity of the moving spins.